X-ray systems are employed for imaging purposes for diagnostic investigation and for interventional operations for example in cardiology, radiology and neurosurgery.
By way of example, FIG. 1 shows such an X-ray system or such an X-ray diagnostic device, which has at least one C-arm 4 mounted rotatably on a stand, on the ends of which are attached an X-ray radiation source 6, for example an X-ray emitter or tube, and an X-ray image detector 5, and a high voltage generator for the generation of tube voltage. The stands holding the C-arm can be fixed to the floor, ceiling or side wall. The C-arm 4 can also be controlled or guided by a robot 8. The X-ray image detector 5 can be a rectangular or square, flat semiconductor detector, which is preferably made of amorphous silicon (aSi).
In the optical path of the X-ray radiation source 6 is located a patient couch 3 to accommodate an area of a patient 7 to be examined. An image system 2 is connected to the X-ray diagnostic device, which receives and processes the image signals from the X-ray image detector 5. The processed image signals can then be displayed on a display device 1 connected to the image system 2.
The X-ray radiation source 6 emits a beam emanating from a beam focus of the X-ray radiation source 6, which strikes the X-ray image detector 5.
The X-ray radiation source 6 and the X-ray image detector 5 in each case travel around the area to be examined, in such a way that the X-ray radiation source 6 and the X-ray image detector 5 lie on opposite sides of the area.
To create 3-D data sets, the rotatably mounted C-arm 4 with the X-ray emitter and X-ray image detector 5 is rotated in such a way that it moves around an area of the patient 7 to be examined (for example the liver) on an orbit of the X-ray radiation source 6 and an orbit of the X-ray image detector 5. The orbits can be traversed completely or partially to produce a 3-D-data set.
X-ray flat detectors are currently being introduced to a large extent universally as X-ray detectors in many areas of X-ray technology, for example in radiography or also in interventional angiography and cardiology.
The ability of today's X-ray flat detectors to make efficient use of the incoming X-ray radiation R for image generation is high, but does not reach the theoretical limit.
FIG. 2 shows an X-ray flat image detector 5, such as is for example used in FIG. 1. The structural principle of such an indirectly converting flat detector, comprising a scintillator, an array comprising photodiodes D and switch element S and an active readout matrix made of amorphous silicon and activation and readout electronics, which has, inter alia, a line driver Z and a multiplexer amplifier M.
Depending on the beam quality, the quantum efficiency for a scintillator SZ made of Cesium Iodide (CsI) with a layer thickness of for example 600 μm is between about 50% and 80%, as is described for example in M. Spahn, “Flat detectors and their clinical applications”, Eur Radiol (2005), 15: 1934-1947. The local frequency dependent DQE(f) (=“Detective Quantum Efficiency”) is hereby subject to an upper limit, and for typical pixel sizes lies for example between 150-200 llm and for local frequencies of 1-2 lp/mm (line pairs per mm) of interest for the applications, even significantly lower.
Other converter materials used in angiography or radiography, such as for example scintillators made of Gd2S20 or direct converters made from Se, exhibit similar basic behavior, lie for currently possible detectors at an even lower level, comparatively speaking. Above the K-edge it is basically the case that with increasingly harder radiation, the absorption and consequently the DQE(f) diminishes.
Harder radiation occurs very frequently in interventional cardiology applications, on the one hand in the case of obese patients, but also as a result of necessary oblique projections, which geometrically call for penetration of a patient of up to 40 cm and more. The task of an X-ray system is also in these cases still to deliver images of acceptable quality. Here a number of effects now overlap to the disadvantage of the image quality:                Current limitations of the X-ray generation system (tubes, generator) result in the fact that in the case of steep angulations the prescribed detector dose can only be achieved at all by increasing the quantum energy. Voltages of 100 kV or more are then required.        hardening of the radiation by means of prefiltering        further hardening of the radiation upon penetration of the patient, as a result of which the contrast is reduced once more.        limitation of the patient's initial dose as in fluoroscopy, which with prescribed detector dose and increasing object thickness (for example steep angulations) can only be achieved through increasing hardening of the radiation.        
However hardened radiation leads—as described above—to a lower quantum efficiency and thus suboptimal image quality.
Some of the most important properties of scintillators are as follows:                High quantum efficiency, in order to make optimum use of the generated dose.        High number of secondary quanta per absorbed X-ray quantum RQ (see FIG. 3) of X-ray radiation R, in order preferably to guarantee a high signal-to-noise gap (“noise” should here be taken to refer to electronic noise).        High optical transparency, in order to ensure as little as possible of the generated light is reabsorbed into the material again,        Narrow point spread function (or high MTF ‘=’ Modulation Transfer Function), that is the property that the generated light can be detected in a locally limited manner perpendicular to the direction of incidence.        
Further properties are for example speed or low lag behavior.
There are various possibilities for realizing subsidiary aspects of the abovementioned properties of scintillators.
In flat detectors conventionally used for angiography and radiography, large-area CsI of needle-shaped structure and 600 μm layer depth or more is employed, for example. The advantages are a good MTF, a high number of secondary quanta and good transparency. The absorption properties are however suboptimal in the case of hard radiation. Since the MTF suffers as a result of greater layer depths (longer transmission paths) and, at the same time, significantly thicker layers are more difficult to manufacture and as the case may be, less economical, this is not a suitable solution.
In computed tomography (CT) Gd2S2O is used, with significantly better absorption properties. The non-structured material nevertheless exhibits considerably worse resolution properties (MTF), which can only be compensated for by means of a discrete structure (optical separation of the scintillators from pixel to pixel).
This can be realized for typical CT-pixel sizes of around 1 mm, but not for pixel sizes of 150-200 μm, such as are used in angiography/radiography. The discrete structure of the scintillator layer in computed tomography additionally permits greater layer thicknesses, so that quantum efficiencies of 95% or more are attained.
FIG. 3 (a)-(d) shows a number of possible embodiments for flat detectors in angiography or radiography. The active readout matrix MX with photodiodes is on the side facing away from the X-ray. The main advantages and disadvantages of these embodiments are listed below:
FIG. 3(a) shows an unstructured scintillator SZ without reflector RF and relatively wide point spread functions PBF (corresponding to a “poor” MTF) through the absence of structuring.
The absorption level influences the PBF (applies to each of the 4 variants), wherein less light yield is produced as a result of loss of light.
FIG. 3(b) shows an unstructured scintillator SZ with reflector RF and still relatively wide point spread functions PBF through an absence of structuring, wherein a higher light yield is generated by the reflector.
FIG. 3(c) shows a structured scintillator SZ: improved PBF through structuring (a part of the light quanta L is directed within the longitudinal crystals (within the total angle of reflection at the CsI-air transition). This solution approach is less economical as a result of the structuring.
FIG. 3(d) shows a structured, thicker scintillator SZ with reflector RF. This results in higher absorption, lower economic viability as a result of structuring and greater thickness, as well as a higher light yield and wider PBF (corresponds to a “poorer” MTF).